The present invention relates to the art of diagnostic nuclear imaging. It finds particular application in conjunction with gamma cameras and single photon emission computed tomography (SPECT), and will be described with particular reference thereto. However, it is to be appreciated that the present invention is also amenable to other like applications.
Diagnostic nuclear imaging, is used to study a radionuclide distribution in a subject. Typically, in SPECT, one or more radiopharmaceuticals or radioisotopes are injected into a subject. The radiopharmaceuticals are commonly injected into the subject""s blood stream for imaging the circulatory system or for imaging specific organs which absorb the injected radiopharmaceuticals. One or more gamma or scintillation camera detector heads, typically including a collimator, are placed adjacent to a surface of the subject to monitor and record emitted radiation. The camera heads typically include a scintillation crystal which produces a flash or scintillation of light each time it is struck by radiation emanating from the radioactive dye in the subject. An array of photomultiplier tubes and associated circuitry produce an output signal which is indicative of the (x, y) position of each scintillation on the crystal. Often, the heads are rotated or indexed around the subject to monitor the emitted radiation from a plurality of directions to obtain a plurality of different views. The monitored radiation data from the plurality of views is reconstructed into a three dimensional (3D) image representation of the radiopharmaceutical distribution within the subject.
One of the problems with this imaging technique is that photon absorption and scatter (i.e., Compton scattering) by portions of the subject between the emitting radionuclide and the camera head distort the resultant image. One solution for compensating for photon attenuation is to assume uniform photon attenuation throughout the subject. That is, the subject is assumed to be completely homogenous in terms of radiation attenuation with no distinction made for bone, soft tissue, lung, etc. This enables attenuation estimates to be made based on the surface contour of the subject. Of course, human subjects do not cause uniform radiation attenuation, especially in the chest.
In order to obtain more accurate radiation attenuation measurements, a direct measurement is made using transmission computed tomography techniques. In this technique, radiation is projected from a radiation source through the subject. The transmission radiation is received by detectors at the opposite side. The source and detectors are rotated to collect transmission data concurrently with the emission data through a multiplicity of angles. This transmission data is reconstructed into an image representation or attenuation map using conventional tomography algorithms. The radiation attenuation properties of the subject from the transmission computed tomography image are used to correct for radiation attenuation in the emission data. See, for example, U.S. Pat. Nos. 5,210,421 and 5,559,335, commonly assigned and incorporated herein by reference.
To assure that the radiation comes along a known path through or from the subject, collimators are often placed in front of radiation-receiving faces of the detector heads. The collimators typically include a grid of lead vanes which assure that received radiation is traveling along a path from the subject substantially perpendicular to the radiation-receiving faces of the detector heads.
Other collimators have been developed to xe2x80x9cmagnifyxe2x80x9d regions of interest. In a cone-beam collimator, the vanes are tapered or angled such that all the vanes point at a common focal point. Radiation reaching the radiation-receiving faces of the detector heads is constrained by the cone-beam collimator to radiation traveling along divergent paths in two directions such that the entire radiation-receiving face of the detector head is used to examine a relatively small region of interest. This magnification improves the resolution in two planar dimensions. Rather than magnifying in two dimensions, fan-beam collimators have also been developed which magnify in one dimension. That is, the vanes are oriented such that the vanes focus the radiation on a focal line, rather than a focal point.
The collimators introduce a system geometric point response that is spatially varying and deteriorates with distance from the face of the collimator. This results in shape distortions and nonuniform density variations in images reconstructed from projection data obtained from a SPECT imaging system. The system geometric point response is dependant on the point source location and collimator geometry. See, for example, G. L. Zeng, et al., xe2x80x9cThree-Dimensional Iterative Reconstruction Algorithms with Attenuation and Geometric Point Response Correction,xe2x80x9d IEEE Trans. Nucl. Sci., Vol. 38, pp. 693-702, 1991.
Scatter correction is an important factor in reconstructing accurate images from SPECT data. Scatter correction techniques, such as multiple-window subtraction and intrinsic modeling with iterative algorithms, have been under study for many years. In fact, methods have been developed for scatter correction in SPECT. See, for example: Z. Liang, et al., xe2x80x9cSimultaneous Compensation for Attenuation, Scatter and Detector response for SPECT Reconstruction in Three Dimensions,xe2x80x9d Phys. Med. Biol., Vol. 37, pp. 587-603, 1992; A. T. Riaku, et al., xe2x80x9cPhoton Propagation and Detection in Single-Photon Emission Computed Tomographyxe2x80x94an Analytic Approach,xe2x80x9d Med. Phys., Vol. 21, p. 1311-1321, 1994; A. T. Riaku, et al., xe2x80x9cExperimental and Numerical Investigation of the 3D SPECT Photon Detection Kernel for Non-Uniform Attenuating Media,xe2x80x9d Phys. Med. Biol., Vol. 41, pp. 1167-1189, 1996; and, R. G. Wells, et al., xe2x80x9cExperimental Validation of an Analytical Method of Calculating SPECT Projection Data,xe2x80x9d IEEE Trans. Nucl. Sci., Vol. 44, pp. 1283-1290, 1997. Methods that use multiple acquisition energy windows to estimate the scattered photons and subtract the estimated photons from the projection data have found applications in research and clinical studies. See, for example: B. Axelsson, et al., xe2x80x9cSubtraction of Compton-Scattered Photons in Single-Photon Emission Computed Tomography,xe2x80x9d J. Nucl. Med., Vol. 25, pp. 490-494, 1984; R. J. Jaszczak, et al., xe2x80x9cImproved SPECT Quantification Using Compensation for Scattered Photons,xe2x80x9d J. Nucl. Med., Vol. 25, pp. 893-900, 1984; K. F. Koral, et al., xe2x80x9cSPECT Dual-Energy Window Compton Correction: Scatter Multiplier Required for Quantification,xe2x80x9d J. Nucl. Med., Vol. 31, pp. 90-98, 1990; K. Ogawa, et al., xe2x80x9cA Practical Method for Position-Dependent Compton-Scatter Correction in SPECT,xe2x80x9d IEEE Trans. Med. Imag., Vol. 10, pp. 408-412, 1991; M. A. King, et al., xe2x80x9cA Dual Photo-Peak Window Method for Scatter Correction,xe2x80x9d J. Nucl. Med., Vol. 33, pp. 605-612, 1992; and, D. R. Gilland, et al., xe2x80x9cA 3D Model of Non-Uniform Attenuation and Detector Response Compensation for Efficient Reconstruction in SPECT,xe2x80x9d Phys. Med. Biol., Vol. 39, pp. 547-561, 1994. These pre-processing methods are efficient and effective, but the pre-subtracting methods tend to increase noise and may introduce negative or zero values at locations where projection values are positive. An alternative method to avoid subtraction is to add estimated scatter events to forward projections of the current reconstructed image in an iterative algorithm. See, for example, J. E. Bowsher, et al., xe2x80x9cBayesian Reconstruction and Use of Anatomical a Priori Information for Emission Tomography,xe2x80x9d IEEE Trans. Med. Imag., Vol. 15, pp. 673-686, 1996. However, subtracting or adding data tends to increase the noise level in the data. Iterative reconstruction methods can model scatter physics in the projector/backprojector and have been shown to provide more accurate reconstructions than subtracting/adding methods. See, for example: C. E. Floyd, et al., xe2x80x9cMaximum Likelihood Reconstruction for SPECT with Monte Carlo Modeling: Asymptotic Behavior,xe2x80x9d IEEE Trans. Nucl. Sci., Vol. 34, pp. 285-287, 1987; J. E. Bowsher, et al., xe2x80x9cTreatment of Compton Scattering in Maximum-Likelihood, Expectation-Maximization Reconstructions of SPECT Images,xe2x80x9d J. Nucl. Med., Vol. 32, pp. 1285-1291, 1991; E. C. Frey, et al., xe2x80x9cModeling the Scatter Response Function in Inhomogeneous Scattering Media for SPECT,xe2x80x9d IEEE Trans. Nucl. Sci., Vol. 41, pp. 1585-1593, 1994; A. Welch, et al., xe2x80x9cA Transmission-Map-Based Scatter Correction Technique for SPECT in Inhomogeneous Media,xe2x80x9d Med. Phys., Vol. 22, pp. 1627-1635, 1995; F. J. Beekman, et al., xe2x80x9cImproved Quantitation in SPECT Imaging Using Fully 3D Iterative Spatially Variant Scatter Compensation,xe2x80x9d IEEE Trans. Med. Imag., Vol. 15, pp. 491-499, 1996; and, F. J. Beekman, et al., xe2x80x9cScatter Compensation Methods in 3D Iterative SPECT Reconstruction: A Simulation Study,xe2x80x9d Phys. Med. Biol., Vol. 42, pp. 1619-1632, 1997. These iterative methods use spatially variant scatter point response functions within the projector/backprojector. Nevertheless, pre-storing a complete set of scatter point response functions for each patient is not feasible in practice. Research has been done to approximate the response functions in an object by the water equivalent depth method. See, for example, E. C. Frey, et al., 1994; and, F. J. Beekman, et al., 1996. However, these approximation methods are still not efficient, and do not work well for non-uniform objects.
It has been demonstrated that a slice-to-slice blurring projector/backprojector is efficient and effective when used in an iterative maximum likelihood expectation maximization (ML-EM) algorithm. See, for example: A. W. McCarthy, et al., xe2x80x9cMaximum Likelihood SPECT in Clinical Computation Times Using Mesh-Connected Parallel Computers,xe2x80x9d IEEE Trans. Med. Imag., Vol. 10, pp. 426-436, 1991; G. L. Zeng, et al., xe2x80x9cIterative Reconstruction of Fluorine-18 SPECT Using Geometric Point Response Correction,xe2x80x9d J. Nucl. Med., Vol. 39, pp. 124-130, 1998; and, J. W. Wallis, et al., xe2x80x9cRapid 3-D Projection in Iterative Reconstruction Using Gaussian Diffusion,xe2x80x9d J. Nucl. Med., Vol 37, p. 63P, 1996. Previous developments include, an efficient slice-to-slice blurring technique to model attenuation and system geometric point response in a projector/backprojector pair. The technique uses an ML-EM algorithm to reconstruct SPECT data. However, it fails to address the scatter problem.
When a patient""s attenuation map is available via a transmission scan, a Compton scatter point response function can be estimated by the Klein-Nishina formula. See, for example: C. H. Tung, et al., xe2x80x9cNon-Uniform Attenuation Correction Using Simultaneous Transmission and Emission Converging Tomography,xe2x80x9d IEEE Trans. Nucl. Sci., Vol. 39, pp. 1134-1143, 1992; and Z. Liang, et al., xe2x80x9cReprojection and Back Projection in SPECT Image Reconstruction,xe2x80x9d Proc. IEEE Enerqy Inform. Technol. Southeast, Vol. EITS-1, pp. 919-926, 1989. A method to estimate the projections of first-order Compton scatter has been developed. See, for example, A. Welch, et al., 1995. This method gives a fairly accurate estimate of the scattered projections from an inhomogeneous media, and has been implemented two-dimensionally (2D). However, the method is limited to a 2D scatter model. It employs a ray or line tracing technique that is too computationally burdensome and time consuming to be useful in clinical applications for 3D scatter correction.
The present invention contemplates a new and improved technique for efficient 3D scatter modeling which overcomes the above-referenced problems and others.
In accordance with one aspect of the present invention, a method of modeling 3D first-order scatter, non-uniform attenuation, and 3D system geometric point response in an ML-EM algorithm to reconstruction SPECT data is provided. It includes performing an initial slice-to-slice blurring operation on a volume of estimated emission source data. The volume of estimated emission source data is represented by a 3D array of voxels. A voxel-by-voxel multiplying of the results from the initial slice-to-slice blurring operation by a volume of attenuation coefficients yields a volume of effective scatter source data. The volume of effective scatter source data is voxel-by-voxel added to the volume of estimated emission source data to produce a volume of combined estimated emission and scatter source data. Finally, a secondary slice-to-slice blurring operation is performed on the volume of combined estimated emission and scatter source data.
In accordance with a more limited aspect of the present invention, performing the initial slice-to-slice blurring includes successively convolving parallel slices of the volume of estimated emission source data with respective blurring kernels, and successively adding previously convolved neighboring slices to their immediately following slices prior to that following slices"" convolution.
In accordance with a more limited aspect of the present invention, after addition of the previously convolved neighboring slice and before its own convolution, slices are point-by-point multiplied by a 2D array of attenuation factors.
In accordance with a more limited aspect of the present invention, the attenuation factors are exponential functions whose exponent is a negative of a linear attenuation coefficient times a distance between respective voxels.
In accordance with a more limited aspect of the present invention, each convolution is implemented as two orthogonal 1D convolutions.
In accordance with a more limited aspect of the present invention, the convolutions approximate a Gaussian scattering probability which is a function of a scattering angle.
In accordance with a more limited aspect of the present invention, the blurring kernels are estimated from a least-squares fit comparison of calculated results to projections from known point sources taken one point source at a time using Monte Carlo simulations.
In accordance with a more limited aspect of the present invention, performing the secondary slice-to-slice blurring includes successively convolving parallel slices of the volume of combined estimated emission and scatter source data with respective blurring kernels, and successively adding previously convolved neighboring slices to their immediately following slices prior to that following slices"" convolution.
In accordance with a more limited aspect of the present invention, the blurring kernels are cross shaped.
In accordance with a more limited aspect of the present invention, the convolutions approximate a system geometric point response function. The system geometric point response function is dependent on physical characteristics of a collimator used during collection of the SPECT data.
In accordance with another aspect of the present invention, a projector/backprojector for use in an image processor employing an EM reconstruction algorithm is provided. It includes a first convolver which successively convolves parallel slices of a volume of image data and adds them to neighboring slices prior to the neighboring slices being convolved. The first convolver employs convolution kernels determined from a Compton scattering function which is a function of a scattering angle. A multiplication processor voxel-by-voxel multiplies a volume of image data from the first convolver by a volume of attenuation coefficients. Again, a second convolver successively convolves parallel slices of a volume of image data from the multiplication processor and adds them to neighboring slices prior to the neighboring slices being convolved. This time, the second convolver employs convolution kernels determined from a system geometric point response function which depends on physical characteristics of a collimator used during collection of measured data.
One advantage of the present invention is its ease of implementation and efficiency.
Another advantage of the present invention is the reduction of imaging artifacts from scattered radiation.
Yet another advantage of the present invention is the reduction of computation time in 3D scatter correction.
Another advantage of the present invention is that it compensates for 3D first-order scatter, non-uniform attenuation, and 3D system geometric point response.
Still further advantages and benefits of the present invention will become apparent to those of ordinary skill in the art upon reading and understanding the following detailed description of the preferred embodiments.